Prosthetic implantable antibacterial surgical mesh

ABSTRACT

The disclosed invention is directed to an implantable surgical prosthetic mesh having a nanofiber comprising one or more antibiotic and a polysaccharide, and a non-polysaccharide polymer deposited on the mesh. The mesh of the invention is shown to be effective in eliminating or minimizing the bacterial population in the mesh and surrounding tissue for at least 14 days from surgical implantation of the mesh.

BACKGROUND OF THE INVENTION Field of the Disclosure

The present invention relates to an implantable surgical mesh having ananofiber comprising one or more antibiotic deposited on the surface ofthe mesh.

Description of Related Art

The “background” description provided herein is for the purpose ofgenerally presenting the context of the disclosure. Work of thepresently named inventors, to the extent it is described in thisbackground section, as well as aspects of the description which may nototherwise qualify as prior art at the time of filing, are neitherexpressly or impliedly admitted as prior art against the presentdisclosure.

A hernia is an abnormal protrusion of intra-abdominal tissue or organthrough a defect in the abdominal wall. Currently, the only effectivetreatment of a hernia is surgical repair of the defect. The use of meshin surgical repair of hernias has dramatically decreased the incidenceof hernia recurrence. However, postoperative mesh infection is a majorcomplication that is difficult to treat. In most cases, surgical removalof the implanted mesh becomes necessary to obviate the complicationwhich may cause a subsequent hernia recurrence and surgicalintervention. Therefore, mesh infection incurs physical, economic andpsychological impact on the patient. The incidence of mesh infection ishigher when a mesh is used in a contaminated surgical field, such as incases of strangulated hernias. Mesh infection is thought to be caused bythe postoperative growth of bacteria present in the surgical field atthe time of mesh insertion. As a foreign body, the mesh is devoid of anyblood supply and, thus, forms a favorable nest for bacterialproliferation.

Nanofibers produced by electrospinning have been shown to be anefficient drug delivery system providing sustained release of drugs[Elakkiya et al. (2012) “Fabrication of electrospun poly l-lactide andcurcumin loaded poly l-lactide nanofibers for drug delivery” Fibers andPolymers, 13: 623-630]. They are capable of delivering medicinesdirectly to internal tissues due to their physical characteristics thatinclude large surface area and permeability combined withinter-connecting pore structure of the fibers leading to efficient drugrelease as well as their ease of fabrication [Khan et al. (2012)“Nanofibers—a new trending nano drug delivery systems” InternationalJournal of Pharmaceutical Research & Analysis. 3: 47-55].

Electrospinning is a promising technique for producing nano-scalefibers, which are excellent candidates for various applications, e.g.,as a wound dressing and delivery system [Li et al. (2004)“Electrospinning of nanofibers: reinventing the wheel” AdvancedMaterials 16: 1151-1170]. It is a relatively simple and adaptableprocess to produce nanofiber from a polymer solution or melt which waspatented by Formhals in the 1930s (U.S. Pat. Nos. 1,975,504, 2,160,962,and 2,187,306—each incorporated herein by reference in their entirety).The method comprises applying a high voltage electrical field betweenthe tip of a nozzle and a collector in order to generate sufficientelectrostatic force to overcome the surface tension of a droplet of thepolymer solution at the nozzle tip. When the surface tension isovercome, the hemispherical surface of the fluid at the tip of thenozzle stretches to form a conical shape known as the Taylor cone[Taylor, G. L. (1969) “Electrically driven jets” Proceedings of theRoyalSociety of London A 313: 453, incorporated herein by reference inits entirety]. Further increase of the electric field's strength willdeform the Taylor cone until a jet stream is extruded from the cone'sapex. During the process, and depending on the solution properties andoperating conditions, the solvent evaporates as the jet moves toward thecollector which decreases the jet radius and increases the polymerconcentration and viscosity. When the solvent is fully evaporated, thejet stretching stops and results in producing fiber of highly reduceddiameter which deposits on the grounded collector in the form of arandom nonwoven structure. The process of the electrospinning is welldescribed, see for example Huang et al. [“A review on polymer nanofibersby electrospinning and their applications in nanocomposites” CompositesScience and Technology (2003) 63(15): 2223-2253; Kim et al.[“Polybenzimidazole nanofiber produced by electrospinning” PolymerEngineering and Science (1999)39 (5): 849-854]; Fang et al. “DNA fibersby electrospinning” Journal of Macromolecular Science (1997) 36 (2):169-173]; and Doshi et al. [“Electrospinning process and applications ofelectrospun fibers” Journal of Electrostatics (1995). 35 (2-3):151-160]—each incorporated herein by reference in their entirety.Nanofibers in the range of 10 to 1000 nm diameter can be achieved bychoosing the appropriate parameters such as viscosity, concentration,applied voltage, distance between the two electrodes, and nozzle tip(needle) diameters. However, the instability, the whipping of the fiber,and bead formation remain important problems in the electrospinningprocess.

Abdel Hady [Abdel-Hady et al. “Experimental validation of upwardelectrospinning process” Nanotechnology, Vol. 2011, 14 pages,http://dx.doi.org/10.5402/2011/851317; and Abdel-Hady et al. InProceedings of the Egyptian Engineers Association Conference, (2009) AlRiyadh, Saudi Arabia—each of which is incorporated herein by referencein its entirety] introduced the upward electrospinning method thatovercomes some of the shortcoming of the previous methods. In themethod, the fiber is produced on a jet directed upward. The combinationof gravitational force and surface tension oppose against theelectrostatic force thereby stretching the fiber.

Abbaspour et al. [Junsishshapur J. Microbiol. (2015) 8 (10): e24239]disclose a chitosan/polyvinyl alcohol and mafenide acetate nanofibercontaining 10% and 30% by weight chitosan and 20 and 40% mafenideacetate. Also, they disclose that films of nanofibers with and withoutmafenide acetate prevent bacterial penetration, but films with mafenideacetate were more effective, in particular against Pseudomonasaeruginosa and Staphylcoccus aureus.

Zupanicic et al. [Mol. Pharmaceutics (2016) 13 (1), 295-305] examine thelong-term sustained release of ciprofloxacin from nanofiber ofhydrophobic polymer such as poly(methyl methylacrylate) (PMMA) andpolycaprolactone (PLC) and a blend of hydrophobic and hydrophilicpolymers such as polyvinyl alcohol (PVA), polyethylene oxide (PEO), andchitosan. PEO containing nanofiber mats display high burst of drugrelease. In contrast, nanofiber containing PVA and chitosan show muchsmall burst. In addition, the reference discloses that the release ofthe drug can be tuned by various blends of PMMA with PVA or chitosan.

Moydeen et al. [Int. J. Biol. Macromol. (2018) 116, 1250-1259] disclosethe fabrication of electrospun polyvinyl alcohol/dextran as a carrier ofdrugs such as ciproflaxin. Also, they examine the release of theciprofloxacin from the nanofiber and show that the core-shell nanofiberscan sustain the ciprofloxacin release compared with the blendingelectrospinning nanofiber.

Murthy et al. [Bull. Mate. Sci. (2017) 40 (4) 645-653] discloseciprofloxacin-loaded chitosan microparticles (CMP) impregnated inchitosan and poly(vinyl alcohol) (PVA) scaffold for effective deliveryof drug in a sustained manner to wound site.

Such conventional drug-impregnated materials are not, however, effectivefor use in surgical treatment of hernias. Conventional materials areunable to provide drug release profiles that provide a steady andpredictable threshold dosage of active material. In additionconventional materials do not exhibit good compatibility with livetissue to promote quick and lasting wound healing, especially for herniasurgery which occurs at a location on a human subject that is exposed tosignificant pressure and pulling strain.

Accordingly one object of the present disclosure is to provide animplantable surgical prosthesis mesh having nanofiber comprising apolysaccharide, one or more antibiotics, and a biocompatible polymerdeposited on the surface of the mesh. The mesh of the invention providescontinues source of antibiotics for more than 14 days to tissue aftersurgical implantation.

The foregoing paragraphs have been provided by way of generalintroduction, and are not intended to limit the scope of the followingclaims. The described embodiments, together with further advantages,will be best understood by reference to the following detaileddescription taken in conjunction with the accompanying drawings.

SUMMARY

A first aspect of the invention is directed to an implantable medicalprosthetic mesh having nanofibers deposited on the surface of a meshsubstrate, wherein the nanofibers comprise a polysaccharide, anantibiotic, and a biocompatible polymer.

In a preferred embodiment, the average diameter of the fiber is in therange of 50 nm to 300 nm and the bore size in the range of 300 nm to 900nm.

In another preferred embodiment, the polysaccharide is chitosan,dextran, cellulose, or combination thereof.

In a more preferred embodiment, the polysaccharide is chitosan.

In another preferred embodiment, the chitosan has a degree ofdeacetylation in the range of 70% to 95%.

In another preferred embodiment, the antibiotic is a penicillin,tetracycline, cephalosporin, quinolone, lincomycin, macrolide,sulfonamide, glycopeptide, aminoglycoside, carbapenem, or combinationthereof.

In a preferred embodiment, the antibiotic is a quinolone antibiotic.

In a preferred embodiment, the antibiotic is selected from the groupconsisting of lomefloxacine, ofloxacin, gatifloxacin, norfloxacin,ciprofloxacin, moxifloxacin, levofloxacin, gemifloxacin, delafloxacin,cinoxacin, nalidixic acid, trovafloxacin, sparfloxacin, and combinationsthereof.

In a particularly preferred embodiment, the antibiotic is ciprofloxacin.

In another preferred embodiment, the polymer is selected from the groupconsisting of polyvinyl alcohol, polyvinylpyralidone, poly(methacrylate)and polycaprolactone.

In another preferred embodiment, the mesh substrate comprisespolypropylene, polytetrafluroethylene, polyethylene terephthalate, orpolyvinylidene fluoride.

In another preferred embodiment, the mesh substrate comprises abiodegradable polymer.

In another preferred embodiment, the mesh is a polyester,polysaccharide, or polyurethane.

In another preferred embodiment, the mesh is a hernia mesh or pelvicmesh.

In another preferred embodiment, the polysaccharide is chitosan, theantibiotic is ciprofloxacin, and the polymer is polyvinyl alcohol orpolyvinylpyrrolidone.

A second aspect of the invention is directed to a method of making animplantable medical prosthetic mesh comprising:

mixing a first solution containing a polysaccharide with a secondsolution comprising an antibiotic and a polymer, and

electrospinning the solution mixture to nanofibers.

In a preferred embodiment, the ratio of the first solution/the secondsolution is in the range of 1 to 5.

In another preferred embodiment, the polysaccharide is chitosan.

In another preferred embodiment, the antibiotic is ciprofloxacin.

In another preferred embodiment, the polymer is polyvinyl alcohol,polyvinylpyrrolidone, or a mixture thereof.

In another preferred embodiment, the electrospinning is carried out bythe upward electrospinning method.

BRIEF DESCRIPTION OF THE DRAWINGS

A more complete appreciation of the invention and many of the attendantadvantages thereof will be readily obtained as the same becomes betterunderstood by reference to the following detailed description whenconsidered in connection with the accompanying drawings, wherein:

FIG. 1 shows a diagram of the upward electrospinning set up.

FIG. 2A shows a photograph and SEM micrographs of sample A atmagnification of ×30000.

FIG. 2B shows a photograph and SEM micrographs of sample B atmagnification of ×30000.

FIG. 3 is a graph illustrating the release percentage ofciprofloxacin-HCl from different formulas in phosphate buffer vs. Timewhere A (Chito:PVP:PVA, 30:50:20%), B (Chito:PVP, 30:70%), and C(Chito:PVA, 30:70%).

FIG. 4 is a graph illustrating DSC thermograms of (A) Ciprofloxacin-HCl,(B) Chitosan, (C) PVP, (D) PVA and (E) physical mixture.

FIG. 5 is a graph illustrating the antibacterial activity ofciprofloxacin released from chitosan/polymer over 14 days incubationperiod at 37° C. of E. coli, S. aureus and P. aeruginosa cultures.

FIG. 6 is a graph illustrating the antibacterial activity ofciprofloxacin released from electrospun nanofiber mats over 14 daysincubation period at 37° C. of E. coli, S. aureus and P. aeruginosacultures.

FIG. 7A is a graph illustrating the antibacterial activity ofciprofloxacin released from the polymer coated electrospun nanofibermats (16 samples) along 14 days incubation period at 37° C. against E.coli.

FIG. 7B is a graph illustrating the antibacterial activity ofciprofloxacin released from the polymer coated electrospun nanofibermats (16 samples) over 14 days incubation period at 37° C. against S.aureus.

FIG. 8 is a graph illustrating the microbial count (CFU/mL) after 14days treatments of 20 rabbits of medicated and non-medicated meshes.

FIG. 9 is a picture illustrating non-medicated mesh.

FIG. 10 is a picture illustrating medicated mesh.

FIG. 11 is a picture illustrating extracted mesh from dorsalsubcutaneous pocket.

DETAILED DESCRIPTION

Embodiments of the present disclosure will now be described more fullyhereinafter with reference to the accompanying drawings, in which some,but not all embodiments of the disclosure are shown. The presentdisclosure will be better understood with reference to the followingdefinitions.

All publications mentioned herein are incorporated herein by referencein full for the purpose of describing and disclosing the methodologies,which are described in the publications, which might be used inconnection with the description herein. Nothing herein is to beconstrued as an admission that the inventors are not entitled toantedate such disclosure by virtue of prior disclosure. Also, the use of“or” means “and/or” unless stated otherwise. Similarly, “comprise,”“comprises,” “comprising” “include,” “includes,” and “including” areinterchangeable and not intended to be limiting.

As used herein, the term “compound” is intended to refer to a chemicalentity, whether in a solid, liquid or gaseous phase, and whether in acrude mixture or purified and isolated.

As used herein, the term “salt” refers to derivatives of the disclosedcompounds, monomers or polymers wherein the parent compound is modifiedby making acid or base salts thereof. Exemplary salts include, but arenot limited to, mineral or organic acid salts of basic groups such asamines, and alkali or organic salts of acidic groups such as carboxylicacids. The salts of the present disclosure can be synthesized from theparent compound that contains a basic or acidic moiety by conventionalchemical methods. Generally such salts can be prepared by reacting thefree acid or base forms of these compounds with a stoichiometric amountof the appropriate base or acid in water or in an organic solvent, or ina mixture of the two; generally non-aqueous media like ether, ethylacetate, ethanol, isopropanol, or acetonitrile are preferred.

As used herein, the term “about” refers to an approximate number within20% of a stated value, preferably within 15% of a stated value, morepreferably within 10% of a stated value, and most preferably within 5%of a stated value. For example, if a stated value is about 8.0, thevalue may vary in the range of 8±1.6, ±1.0, ±0.8, ±0.5, ±0.4, ±0.3,±0.2, or ±0.1.

As used herein a “polymer” or “polymeric resin” refers to a largemolecule or macromolecule, of many repeating subunits and/or substancescomposed of macromolecules. As used herein a “monomer” refers to amolecule or compound that may bind chemically to other molecules to forma polymer. As used herein the term “repeat unit” or “repeating unit”refers to a part of the polymer or resin whose repetition would producethe complete polymer chain (excluding the end groups) by linking therepeating units together successively along the chain. The method bywhich monomers combine end to end to form a polymer is referred toherein as “polymerization” or “polycondensation”, monomers are moleculeswhich can undergo polymerization, thereby contributing constitutionalrepeating units to the structures of a macromolecule or polymer. As usedherein “resin” or “polymeric resin” refers to a solid or highly viscoussubstance or polymeric macromolecule containing polymers, preferablywith reactive groups. As used herein a “copolymer” refers to a polymerderived from more than one species of monomer and are obtained by“copolymerization” of more than one species of monomer. Copolymersobtained by copolymerization of two monomer species may be termedbipolymers, those obtained from three monomers may be termed terpolymersand those obtained from four monomers may be termed quarterpolymers,etc. As used herein, “cross-linking”, “cross-linked” or a “cross-link”refers to polymers and resins containing branches that connect polymerchains via bonds that link one polymer chain to another. The polymer maybe natural or synthetic. Natural polymer includes, but not limited topolysaccharides such as, but not limited to cellulose, dextran, andnatural polyesters. Synthetic polymers include but not limited topolypropylene, polyvinylalcohol, polyvinylpyralidone,polytetrafluroethylene, polyethylene terephthalate, and polyvinylidenefluoride.

A first aspect of the invention is directed to an implantable medicalprosthetic mesh having nanofibers deposited on the surface of a meshsubstrate, wherein the nanofibers comprise:

a polysaccharide in an amount in the range of 8 wt. % to 35 wt. %,

an antibiotic in an amount in the range of 15 wt. % to 35 wt. %, and

a polymer in an amount in the range of 45 wt. % to 70 wt. %.

Surgical mesh is a loosely woven or knitted sheet which is used as apermanent or temporary support for organs and other tissues duringsurgery. Surgical mesh is created from both inorganic and biologicalmaterials and is used in a variety of surgeries. Permanent meshes remainin the body, whereas temporary ones dissolve over time.

In some embodiments, yarns include at least two filaments which may bearranged to create openings there between, the yarns also being arrangedrelative to each other to form openings in the mesh. Alternatively, themesh may be formed from a continuous yarn that is arranged in loops thatgive rise to the openings in the mesh. The use of a mesh having yarnsspaced apart in accordance with the present disclosure has the advantageof reducing the foreign body mass that is implanted in the body, whilemaintaining sufficient tensile strength to securely support the defectand tissue being repaired by the mesh. Moreover, the openings of themesh of the present disclosure may be sized to permit fibroblastthrough-growth and ordered collagen laydown, resulting in integration ofthe mesh into the body. Thus, the spacing between the yarns may varydepending on the surgical application and desired implantcharacteristics as envisioned by those skilled in the art. Moreover, dueto the variety of sizes of defects, and of the various fascia that mayneed repair, the mesh may be of any suitable size.

In embodiments in which at least two filaments form a yarn, thefilaments may be drawn, oriented, crinkled, twisted, braided, commingledor air entangled to form the yarn. The resulting yarns may be braided,twisted, aligned, fused, or otherwise joined to form a variety ofdifferent mesh shapes. The yarns may be woven, knitted, interlaced,braided, or formed into a surgical mesh by non-woven techniques. Thestructure of the mesh will vary depending upon the assembling techniqueutilized to form the mesh, as well as other factors, such as the type offibers used, the tension at which the yarns are held, and the mechanicalproperties required of the mesh.

In some embodiments, knitting may be utilized to form a mesh of thepresent disclosure. Knitting involves, in embodiments, the intermeshingof yarns to form loops or inter-looping of the yarns. In embodiments,yarns may be warp-knitted thereby creating vertical interlocking loopchains, and/or yarns may be weft-knitted thereby creating rows ofinterlocking loop stitches across the mesh. In other embodiments,weaving may be utilized to form a mesh of the present disclosure.Weaving may include, in embodiments, the intersection of two sets ofstraight yarns, warp and weft, which cross and interweave at rightangles to each other, or the interlacing of two yarns at right angles toeach other. In some embodiments, the yarns may be arranged to form a netmesh which has isotropic or near isotropic tensile strength andelasticity.

In some embodiments, the yarns may be nonwoven and formed bymechanically, chemically, or thermally bonding the yarns into a sheet orweb in a random or systematic arrangement. For example, yarns may bemechanically bound by entangling the yarns to form the mesh by meansother than knitting or weaving, such as matting, pressing,stitch-bonding, needlepunching, or otherwise interlocking the yarns toform a binderless network. In other embodiments, the yarns of the meshmay be chemically bound by use of an adhesive such as a hot meltadhesive, or thermally bound by applying a binder such as a powder,paste, or melt, and melting the binder on the sheet or web of yarns.

The mesh may be a composite of layers, including a fibrous layer asdescribed above, as well as porous and/or non-porous layers of fibers,foams, and/or films. A non-porous layer may retard or prevent tissueingrowth from surrounding tissues, thereby acting as an adhesion barrierand preventing the formation of unwanted scar tissue. In embodiments, areinforcement member may be included in the composite mesh. Suitablemeshes, for example, include a collagen composite mesh such as PARIETEX™(Tyco Healthcare Group LP, d/b/a Covidien, North Haven, Conn.).PARIETEX™ composite mesh is a 3-dimensional polyester weave with aresorbable collagen film bonded on one side. Examples of other mesheswhich may be utilized include those disclosed in U.S. Pat. Nos.6,596,002; 6,408,656; 7,021,086; 6,971,252; 6,695,855; 6,451,032;6,443,964; 6,478,727; 6,391,060; and U.S. Patent Application PublicationNo. 2007/0032805, the entire disclosures of each of which areincorporated by reference herein.

The primary function of surgical mesh is to support prolapsed organseither temporarily or permanently. It is most commonly used in herniasurgery within the abdomen, which is required when an organ protrudesthrough abdominal muscles. Also, surgical mesh may be used for pelvic orvaginal wall reconstructions in women and is implemented to act as agrowth guide for damaged tissue. Ideally, these implants should bestrong enough to survive mechanical loads and actions of whichever bodyarea they become a part of.

Mesh implantation is accompanied by a physiological response to theinserted mesh. A mesh constructed from biocompatible material wouldproduce desired minimal response such as formation of fibrosis aroundthe prosthesis, i.e., scar tissue formation. In some cases, an acuteinflammatory reaction may be triggered due to the formation of giantcells and subsequently granulomas indicating that the tissue is“tolerating” the mesh material. In some other cases, a severeinflammatory response is developed during the integration of the meshinto the tissue leading to fibroblastic cell proliferation. Ultimately,the goal for surgical mesh development is to formulate one that has aminimal in vivo reaction to maximize comfort for the patient, avoidinfection, and ensure smooth integration into the body for tissue withminimal response of the body to a foreign object.

A number of factors are involved in mesh biocompatibility. Mesh porosityis the ratio of pore to total area that plays a role in the developmentof bacterial infection or smooth tissue regeneration. Pore sizes below10 μm are susceptible to infection because bacteria may enter andproliferate, whereas macrophages and neutrophils are too large to passthrough and aid in fighting the bacteria. At pore sizes larger than 75μm, fibroblasts, blood vessels, and collagen fibers infiltrate the meshas part of tissue regeneration. Although there is no general consensuson the best pore size, it can be deduced that larger pores are betterfor development of tissue and integration in vivo.

Any biocompatible polymer may be utilized in formation of a meshsubstrate of the invention. The selected polymer for the mesh substratefor the invention is dependent on the intended use of the mesh. Thepolymer can be selected from a wide array of natural and syntheticbiocompatible polymers that include but is not limited to polypropylene,polyethylene terephthalate, polytetrafluroethylene polyvinylidenefluoride (PVDF), polylactic acid, polyglycolic acid, cellulose,dextrose, or combination thereof. In particular, PVDF mesh is resistantto hydrolysis and sufficiently flexible. Its use is described for bothhernia and pelvic/vaginal wall surgery and is produced via fiberplacement layer by layer.

Synthetic meshes are grouped as heavyweight or lightweight. The weightof the mesh depends on both the weight of the polymer and the amount ofmaterial used. Heavyweight meshes use thick polymers, have small poresize and high tensile strength. The weigh typically about 100 g/cm² (1.5g for 10×15 cm mesh). In contrast, lightweight meshes are composed ofthinner filaments with relatively large pores (>1 mm), and weightypically about 33 g/cm² (0.5 g for 10×15 cm mesh). They contain lessmaterial, initiate less pronounced foreign body reaction, and are moreelastic.

A deterioration of the tensile strength of the mesh or an increase inthe ability of the mesh material to stretch could potentially lead tohernia recurrence or a poor functional result. Therefore, mesh materialsmust also possess the biomechanical properties necessary to withstandthe stresses placed on the abdominal wall. This means that once thesurgical mesh is implanted, some of the flexibility of the abdominalwall should be preserved. The natural elasticity of abdominal wall at 32N/cm is about 38%. Lightweight meshes have an elasticity of about 20-35%at 16 N/cm, whereas heavyweight meshes have half of this elasticity(4-15% at 16 N/cm), and can restrict abdominal distension. Therefore,choosing a stronger mesh prosthetic should proceed cautiously. On theother hand, strain values greater than 30% indicate that these materialsmay stretch more than the native human abdominal wall, so may notmaintain a functional repair and could result in bulging or recurrence.

Strength depends on filament type (multi- or mono-), woven or knitted,and the polymer type. Knitted meshes have greater flexibility and largerpores, but are not as strong compared to woven meshes. Knitted meshescan be stretched in any direction, whereas woven meshes allow stretchingonly in the direction oblique to the ninety degree intersection of thetheir strands. Orientation of the mesh also affects physico-mechanicalproperties during implantation. Therefore, meshes with anisotropicstretchability should be orientated with the most stretchable axis inthe direction of least overlap to prevent early mesh dislocation.

The maximum intra-abdominal pressure generated in healthy adults occursduring jumping, and estimated to be about 170 mmHg. Meshes used torepair large hernias need to withstand at least 180 mmHg beforebursting. All synthetic meshes are sufficiently strong. Most commonlyused mesh prosthetics have a tensile strength of at least 32 N/cm. Thisis easily achieved as even the lightest meshes will withstand twice thispressure without bursting (burst pressure of Vypro 360 mmHg). This meansthat the tensile strengths of more than 100 N/cm of conventionalheavyweight meshes (e.g. Prolene) are disproportional and not necessaryfor an effective repair. Therefore, mesh hernia repair failure oftenresults from the separation of the mesh-fascia interface; not the meshfailing.

Composite meshes combine more than one material. The main advantage of acomposite mesh is that it can be used in the intraperitoneal space withminimal adhesion formation. They require a specific orientation: thevisceral side has a micro porous surface to prevent visceral adhesions,whereas the non-visceral side is often macro porous to allow parietaltissue ingrowth. The meshes are preferably one or other ofpolypropylene, polyester, and polytetrafluoroethylene, which are used incombination with each other or with additional materials such astitanium, omega 3, monocryl, polyvinylidene fluoride (PVDF) andhyaluronate.

There are two categories of composite meshes: absorbable and permanent.Barrier coatings in absorbable composite meshes require hydration priorto usage, and they are not amenable to modification, so they cannot becut. However, they allow for neoepithelialization of the mesh beforevisceral adhesion, which mitigates viscera-mesh related complications,and can aid in tissue ingrowth. Parietex® composite mesh was the firstto offer a resorbable collagen barrier on one side to limit visceralattachments and a three-dimensional polyester knit structure on theother to promote tissue ingrowth and ease of use. The collagen film iscomposed of glycerol, polyethylene glycol, and porcine collagen. Thisbalance of material properties produces superior cellular proliferationwhen compared to polypropylene mesh in vitro and works with the body'snatural systems to provide rapid fibrous ingrowth, minimal shrinkage,and strong tissue integration.

Permanently combined meshes take advantage of the properties of bothmacro and micro porous meshes. A micro porous mesh permits placementadjacent to viscera, whereas macro porous mesh promotes parietal tissueingrowth. These meshes can be modified and are easily cut to fitspecific applications. These properties permit intraperitoneal placement(e.g. Dual Mesh®, Dulex®, and Composix®).

Since the intended use of a mesh is to repair and support organs due todefects, the mesh may be configured to any suitable shape or size thatis conducive to facilitating the correction or repair of the particulardefect. The prosthetic mesh of the present invention may be manufacturedby conventional techniques. For example the mesh may be cut into anappropriate shape by a cutting blade, hot knife, or ultrasonic cuttingdevice to any size or shape appropriate for the intended use. The threedimensional structure of the device can be formed by shaping a sheet ofmesh over a heated tool and cooling the sheet material to set the shape.In the case of biocompatible synthetic polymers such as but not limitedto polypropylene knitted mesh, the mesh may be thermoformed over aheated tool in the range of from about 80° C. to about 160° C. andpreferably about 118° C. under suitable pressure for a suitable amountof time in the range of 30 s to 30 min. The thermoforming should beconducted in a manner that does not distort or weaken the meshstructure. Generally polypropylene mesh should be compressed between amale and female forming tool which are closed together at a speed in therange of from about 2 to about 25 cm/min and preferable at a speed of inthe range of from about 5 to about 7.5 cm/min. Polypropylene meshesshould generally be heated for in the range of from about 2 to about 5min and cooled to below room temperature to set the shape. Currently, itis preferred to cool polypropylene mesh to about 12° C. or less for atime in the range of 2 to about 5 min.

Several natural polysaccharides may be used to make the nanofibers to bedeposited on the mesh of the invention. They include, but are notlimited to hemicelluloses, cellulose, chitosan, and pectin. A singlepolysaccharide in the nanofibers or a combination of polysaccharide maybe used in the nanofiber. Since polysaccharides biodegrade over time,the selection of the polysaccharide may affect the rate of release ofthe antibiotic from the mesh of the invention. In a preferredembodiment, the polysaccharide used in the nanofiber is chitosan. Theamount of the polysaccharide constituent of the nanofiber is in therange of 2 wt. % to 55 wt. %, preferably in the range of 4 wt. % to 45wt. %, more preferably 6 wt. % to 40, and most preferably in the rangeof 8 wt. % to 35 wt. % of the total weight of the nanofibers. In aparticularly preferred embodiment, the polysaccharide is present in anamount in the range of 10 wt. % to 26 wt. % of the total weight of thenanofibers.

As used herein, chitosan is intended to refer to a chitin-derivedpolymer that is at least 50% deacetylated, preferably at least 70%deacylated, more preferably at 80% deacetylated, and most preferably atleast 85% deacetylated. It is a linear polysaccharide composed ofrandomly distributed β-(1-4)-2-amino-2-D-glucosamine andβ-(1-4)-2-acetamido-2-D-glucoseamine (acetylated) units, and derivedfrom chitin, a naturally occurring polymer. Chitin is a white, hard,inelastic, nitrogenous polysaccharide isolated from fungi, mollusks, orfrom the exoskeletons of arthropods such as crustaceans and insects. Themajor procedure for obtaining chitosan is the alkaline deacetylation ofchitin with strong alkaline solution. Generally, the raw material iscrushed, washed with water or detergent, and ground into small pieces.After grinding, the raw material is treated with alkali and acid toisolate the polymer from the raw crushed material. The polymer is thendeacetylated by treatment with alkali. Chitin and chitosan differ intheir degrees of deacetylation (DDA). Chitin has a degree ofdeacetylation of 0%, whereas pure chitosan has a degree of deacetylationof 100%. Typically, when the degree of deacetylation is greater thanabout 50% the polymer is referred to as chitosan. Chitosan is a cationicweak base that is substantially insoluble in water and organic solvents.Typically, chitosan is fairly soluble in dilute acid solutions, such asacetic, citric, oxalic, propionic, ascorbic, hydrochloric, formic, andlactic acids, as well as other organic and inorganic acids. Chitosan'scharge gives it bioadhesive properties that allow it to bind tonegatively charged surfaces, such as biological tissues present at asite of trauma or negatively charged implanted devices. Chitosan'sdegree of deacetylation affects it resorption. Chitosan compositionshaving a 50% degree of deacetylation are highly degradable in vivo. Asthe degree of deacetylation increases, chitosan becomes increasinglyresistant to degradation. Chitosan compositions having a degree ofdeacetylation that is higher than 95% degrade slowly over weeks ormonths. In the body chitosan is degraded by lysozyme,N-acetyl-O-glucosaminidase and lipases. Lysozyme degrades chitosan bycleaving the glycosidic bonds between the repeating polysaccharideunits. The byproducts of chitosan degradation are oligosaccharides andglucosamine that are gradually absorbed by the human body. Therefore,when chitosan is used for the local delivery of therapeutic orprophylactic agents, no secondary removal operation is required.

The chitosan used in the nanofibers to be deposited on the implantablesurgical mesh has a degree of deacetylation in the range of 50% to 95%,preferably in the range of 60% to 90%, more preferably in the range of75% to 85% and most preferably about 85%. In some embodiment, thechitosan used has a uniform degree of deacylation. As used herein, theterm “uniform degree of deacetylation” refers to a chitosan compositionmade from a single type of chitosan such as, but not limited to 61DDA,71DDA, or 81DDA. Thus, a composition having a uniform degree ofdeacetylation excludes chitosan compositions having chitosans havingdifferent degrees of deacetylation. In some other embodiments, it may bedesirable to use a composition comprising a mixture chitosan ofpreparation having different degrees of deacetylation.

There are many antibiotics known in the art of which one or more may beincorporated into the nanofibers. Some of the antibiotics display abroad spectrum bactericidal activity; whereas others have a narrowspectrum of bactericidal activity. In some locations such as hospitals,antibiotic resistant strains may be developed due to common use of aspecific antibiotics or class of antibiotics. In such a case, a cocktailof antibiotics from one or more classes of antibiotics may be used inthe nanofibers. The amount of the total antibiotic in the nanofiber isin the range of 5 wt. % to 55 wt. %, preferably in the range of 10 wt. %to 45 wt. %, more preferably 15 wt. % to 40, and most preferably in therange of 20 wt. % to 30 wt. % of the total weight of the nanofibers. Ina particularly preferred embodiment, the total antibiotic is in anamount in the range of 20 wt. % to 25 wt. % of the total weight of thenanofibers. The antibiotics are divided into ten classes:

-   -   (1) Penicillins including but not limited to, Amoxicillin,        Ampicillin, Carbenicillin, Piperacillin, Ticarcillin, Penicillin        g benzathine, Procaine penicillin, and Penicillin v potassium,        Oxacillin, Dicloxacillin, and Nafcillin. Since beta-lactamase is        an enzyme which many bacteria acquire to become resistant to        penicillin, a beta-lactmase inhibitor may be used with some of        the penicillins to obviate antibiotic resistance, in particular,        with penicillins such as clavulanate, sulbactam, and tazobactam.    -   (2) Tetracyclines including, but not limited to, Tetracycline,        Doxycycline, Demeclocycline, Minocycline, Oxytetracycline,        Omadacycline, Sarecycline, and Eravacycline    -   (3) Cephalosporins including, but not limited to, Cefadroxil,        Cephradine, Cefazolin, Cefazolin, Cephalexin, Cefepime,        Ceftaroline, Cefoxitin, Loracarbef, Cefprozil, Cefuroxime,        Cefotetan, Loracabef, Ceftibuten, Ceftriaxone, Cefotamime,        Cefdinir, Cefixime, Cefdinir, Cefditoren, and Ceftazidime,    -   (4) Quinolones including, but not limited to, Lomefloxacine,        Ofloxacin, Gatifloxacin, Norfloxacin, Ciprofloxacin,        Moxifloxacin, Levofloxacin, Gemifloxacin, Delafloxacin,        Cinoxacin, Nalidixic acid, Trovafloxacin, and Sparfloxacin.    -   (5) Lincomycins including, but not limited to, Lincomycin and        Clindamycin.    -   (6) Macrolides include, but not limited to, Telithromycin,        Erythromycin, Azithromycin, Clarithromycin, and Fidaxomicin.    -   (7) Sulfonamides including, but not limited to, Sulfisoxazole        and Sulfamethoxazole.    -   (8) Glycopeptides including, but not limited to, Vancomycin,        Oritavancin, and Telavancin.    -   (9) Aminoglycosides including, but not limited to, Paromomycin,        Tobramycin, Gentamicin, Amikacin, Kanamycin, Neomycin, and        Plazomycin.    -   (10) Carbapenems including, but not limited to, Ertapenem,        Cilastatin, and Imipenem.

In some preferred embodiments, quinolone antibiotics are preferablyincorporated alone or in combination with other antibiotics because oftheir broad spectrum bactericidal activity. In a particular preferredembodiment, ciprofloxacin or salt thereof is used as the antibiotic usedin the nanofibers deposited on the surgical mesh.

The polymer of the nanofiber may be the same or different from that ofthe mesh substrate. It can be natural or synthetic as long as it is abiocompatible polymer. Natural biocompatible polymers include, but arenot limited to, polysaccharides such as, but not limited to cellulose,hemicellulose, pectin, dextran, and natural polyesters such as, but notlimited to, polyglycolic acid or polylactic acid. Synthetic polymersinclude but not limited to polypropylene, polyurethane, polycarbonate,polyvinyl chloride, poly(methyl methacrylate), polyvinyl alcohol,polyvinylfluoride, polyvinylpyrrolidone, polytetrafluroethylene,polyethylene terephthalate, polyvinyledene fluoride, and the like. Insome embodiments, it may be desirable to use one polymers having desiredcharacteristics, whereas in other embodiments an engineered blend ofpolymer may be used to obtain the desired characteristics. The amount ofthe total polymer in the nanofiber is at least 20 wt. %, preferably atleast 30 wt. %, more preferably at least 40 wt. %, and most preferablyat least 50 wt. % of the total weight of the nanofibers. In aparticularly preferred embodiment, the total antibiotic is in an amountin the range of 50 wt. % to 60 wt. % of the total weight of thenanofibers.

The nanofibers deposited on the mesh substrate have an average diameterin the range of 5 nm to 900 nm, preferably in the range of 15 nm to 700nm, more preferably in the range of 30 nm to 550 nm, even morepreferably in the range of 40 nm to 400 nm, and most preferably in therange of 50 nm to 300 nm. Also, they have a bore size in the range of 50nm to 1000 nm, preferably in the range 200 nm to 900 nm, more preferablyin the range of 300 nm to 800 nm, and most preferably in the range of500 nm to 700 nm. The amount of the nanofibers deposited on the mesh isin the range of 1.0 wt % to 30 wt %, preferably in the range of 5 wt %to 20 wt %, preferably 10 wt. % to 15 wt % of the total weight of themesh and the nanofibers.

The nanofiber deposited on the implantable surgical mesh functions as along term antibiotic release system. As used herein, the term “long termantibiotic release” means that the antibiotic is eluted steadily fromthe fiber in a time frame in the range of one day to thirty days,preferably in the range of 5 days to 25 days, more preferably in therange 10 days to 20 days, and most preferably in the range of 13 days to16 days. In a particularly preferred embodiment, the antibiotic isreleased from the fiber steadily for at least 14 days. Generally, thepolymer has a drug release rate in the range of about 0.001 μg/cm² minto about 100 μg/cm² min, preferably in the range of about 0.01 μg/cm²min to 10 μg/cm² min.

The medicated nanofibers may be incorporated into the mesh substrate byseveral methods including but not limited to incorporating the nanofiberin the polymer threads to woven or knit the mesh, threading thenanofibers through the pore of a mesh substrate, and gluing or weldingthe nanofiber on the surface of the mesh substrate using a biocompatibleadhesives. Several known non-toxic adhesives well-known in the art suchas but not limited to collagen-based adhesive, a plant based adhesivessuch as but not limited to Arabic gum, Canada balsam, latex, and starch,and synthetic biocompatible adhesives. For example, several epoxyadhesives are commercially available for use in medical devices such asbut not limited to USP class VI and ISO 10933 are both certified forbiocompatibility. Ultrasound welding is a well-known method of weldingpolymers using ultrasound energy. The nanofibers can be deposited on thesurface of a mesh substrate and welded by ultrasound. The method doesnot require any heat or adhesives.

A second aspect of the invention is directed to a method of making thenanofiber to be deposited on the implantable surgical mesh. The methodcomprises:

mixing a first solution containing a polysaccharide in an amount in therange of 0.5 wt. % to 5 wt. % with a second solution comprising anantibiotic in an amount in the range of 1.0 wt. % to 5.0 wt. % and apolymer in an amount in the range of 6 wt. % to 15 wt. %, and

electrospinning the solution mixture to form the nanofibers.

The solutions may be mixed in any ratio that results in nanoparticlescapable of releasing the antibiotic in a desired time frame for theintend application. The ratio of the first solution/the second solutionis in the range of 1 to 20, preferably in the range of 1 to 15, morepreferably in the range of 1 to 10, even more preferably 1 to 5, andmost preferably 1 to 1.

Electrospinning refers to a process that generates fine polymer fibersusing an electrical charge, typically on the micro or nano scale, fromsolutions comprising polymer or a mixture of polymers. The process doesnot require the use of coagulation chemistry or high temperatures toproduce solid threads from solution, which makes the processparticularly suited to the production of fibers using large and complexmolecules. In the instant invention, the electrospinning methodgenerates fibers from a solution comprising a polysaccharide, anantibiotic, and a biocompatible polymer. When a sufficiently highvoltage is applied to a liquid droplet, the body of the liquid becomescharged, and electrostatic repulsion counteracts the surface tension andthe droplet is stretched; at a critical point a stream of liquid eruptsfrom the surface. The point of eruption is known as the Taylor cone. Ifthe molecular cohesion of the liquid is sufficiently high, streambreakup does not occur and a charged liquid jet is formed. As the jetdries in flight, the mode of current flow changes from ohmic toconvective as the charge migrates to the surface of the fiber. The jetis then elongated by a whipping process caused by electrostaticrepulsion initiated at small bends in the fiber, until it is finallydeposited on the grounded collector. The elongation and thinning of thefiber resulting from bending instability lead to the formation ofuniform fibers with nanometer-scale diameters.

Although there are several architectures for electrospinning apparatuswhich are used in the process to obtain nanofibers, they all sharecommon features. For example, a standard laboratory setup forelectrospinning consists of a spinneret, typically a hypodermic syringeneedle, connected to a high-voltage in the range of 5 to 50 kV of adirect current power supply, a syringe pump, and a grounded collector. Apolymer solution, sol-gel, particulate suspension or melt is loaded intothe syringe and is extruded from the needle tip at a constant rate by asyringe pump. Alternatively, the polymer solution may be fed to theneedle from a tank at constant pressure. The constant pressure feed-typeworks better for lower viscosity feedstocks. Any electrospinningapparatus may be used to prepare the nanofibers of the invention. Insome preferred embodiments, the upward electrospinning method describedherein and Abdel Hady [Abdel-Hady et al. Nanotechnology, Vol. 2011, 14pages, http:—//dx.doi.org/10.5402/2011/851317—incorporated herein byreference in its entirety] is used to obtain the nanofibers of theinvention. The upward electrospinning method utilizes a set up diagramedin FIG. 1. It is shown to be effective in producing nanofibers in therange of 50 nm to 1000 nm, and has the advantage of eliminating thewhipping instabilities; and in turn produces a well-definednanostructure. The concentration of the polymer solution has a majoreffect on the diameter of the nanofiber. Increasing the concentrationleads to increasing in fiber diameter. Another factor affecting thefiber diameter is the distance between the needle and collector.Increasing the distance between the needle and collector leads to adecreasing in the nanofibers diameter. The change in the applied highvoltage on nanofibers diameter has a linear relationship; as the appliedvoltage increases, the nanofibers diameter decreases. Increasing thefeed rate leads to increasing the fiber diameter and bead formation.Smaller needle diameter yields fibers with smaller diameter; yet pumpinga viscous liquid through a needle of small internal diameter may notalways be practical.

The implantable surgical prosthesis mesh of the invention is intendedfor use in any surgical procedure that requires tissue or organ supportto minimize postsurgical complication from infection.

Hernia surgery is one of the most common current applications ofsurgical mesh. Hernias occur when organs or fatty tissue bulge throughopenings or debilitated areas of muscle, usually in the abdominal wall.Surgical mesh is implanted to strengthen tissue repair and minimize therate of recurrence. The surgery can be performed laparoscopically, i.e.,internally, or open with a variety of materials available forprosthesis. Polypropylene is one of the most frequently used types ofmesh, although it may be uncomfortable for the patient afterimplantation. Polyethylene terephthalate (PET) is less utilized inhernia surgery but causes complications with time due to biodegradationin vivo after few years of implantation that nullifies the effects ofthe surgery. Polytetrafluorethylene (PTFE) is used as well, but ismanufactured in the form of a foil and has difficulty integrating intosurrounding tissues.

Similar to hernia surgery, synthetic meshes may be used for organprolapses in the pelvic region as well. Pelvic organ prolapse occurs in50% of women above the age of 50 with a history of one or more vaginalchildbirths throughout her lifetime. Mesh surgery can be performed invarious areas of the pelvic region, such as cystocele, rectocele, andvaginal vault or uterus. While the most commonly used material ispolypropylene, which has acceptable biocompatibility within the pelvicregion, other polymers may be utilized in construct such a surgicalmesh. The vaginal wall has three layers: tunica mucosa, muscularis, andadventitia. When prolapse occurs, smooth fibers of the muscularis arecompromised leading in many cases to increase stiffness in the pelvisarea, in particular, among post-menopausal women. Surgical mesh that isused in pelvic reconstruction must counter this stiffness, but if themodulus of elasticity is too high, it will not sufficiently support theorgans. On the other hand, if the mesh is too stiff, tissue will erodeand inflammatory responses will cause post-surgical complications.

Example 1 Materials:

Unless otherwise specified, all chemicals including 85% degreeN-deacetylated chitosan, PVP (polyvinypyrrolidone), PVA (polyvinylalcohol), glacial acetic acid, potassium dihydrogen phosphate, potassiummonohydrogen phosphate, orthophosphoric acid, ciprofloxacin (CP) andother reagents were fine grades purchased from Sigma Aldrich, USA.Ciprofloxacin-HCl was supplied from Fluka Biochemical, Germany.

Staphylococcus aureus (ATCC 12228), Escherichia coli (ATCC 25922), andPseudomonas aeruginosa (ATCC 27853) were used for evaluation of theantibacterial activity of polymer coated electrospun nanofiber matsamples.

Methods: Polymer Solutions:

Table 1 summarizes the compositions of the polymer solutions containingciprofloxacin-HCl to prepare the corresponding electrospinning polymersolution. A 2 wt % solution of chitosan in 2% (v/v) aqueous acetic acidwas mixed with aqueous solution of 10 wt. % of PVP and/or PVA comprising10 wt. % of ciprofloxacin/HCl.

TABLE 1 Composition of nanofiber solution. Polymer solution Formula AFormula B Formula C Chitosan 2% 30% 30% 30% ^(a)PVP 10% 50% 70% —^(a)PVA 10% 20% — 70% ^(a)comprising 10 wt. % ciprofloxacin-HCl

Preparation of Medicated and Unmedicated Films:

A chitosan (CH) solution was prepared by dissolving 1.5 g CH in 100 mL2% acetic acid by stirring overnight at room temperature. A 5 mL of thesolution was poured into circular Teflon mold (7.3 cm diameter and 1 mmdepth). The mold was covered with an inverted funnel to control solventevaporation at 40° C. for 24 h, then the inverted funnel was removed andleft the mold at the same temperature for another 24 h. The dried filmwas then transferred to a desiccator containing silica gel for 24 hbefore test.

For preparation of medicated cast, a ciprofloxacin (CP) solution (2.5%(wt./v) was prepared in 2% acetic acid using a sonicator (ElmaSchmidbauer GmbH, Switzerland) to solubilize the drug. Then, the CPsolution was mixed with the CH solution described) above in a ratio of1:5. The mixed solution was casted as described above. The finalconcentration of CP in the cast is approximately 33.5% (W/W) in CH.

Example 2 Preparation of Nanofibers by Electrospinning

A custom design of an upward electrospinning apparatus diagramed in FIG.1 was manufactured by Shenzhen Tong Li Tech, China. A set of experimentshave been carried out in order to determine suitable system parameterssuch as the high voltage, distance between the two electrodes, andneedle diameter as well as the solution parameters including percentagesof the different component in the solution, viscosity, and surfacetension to identify suitable solution for spinning.

Medicated and non-medicated CH solutions were prepared as describedabove. But, the concentration of CP was reduced to 0.25% (wt./v) in 2%acetic acid solution (final concentration of CP in the nanofiber isapproximately 3.35 wt. % in CH. The upward spinning method was used forproduction of the nanofibers from CH medicated or non-medicatedsolutions. The solutions were loaded into a glass syringe with a 28gauge needle. The needle had outer diameter of 0.64 mm and innerdiameter of 0.34 mm. The syringe was mounted on a syringe pump which wasadjusted to deliver 2 mL/hour. The pump comprising the syringe waspositioned vertically under the collector inside the machine chamber(see FIG. 1). The distance between the tip of the needle and thecollector surface was adjusted to be 15 cm. The collector is a cylinderrotating and translating horizontally to obtain quasi uniformdistribution of the fiber. The cylinder was covered with an aluminumfoil to collect the fibers. The positive electrode of the power supplywas connected to the collector metal cylinder, whereas the negativeelectrode was connected to the metal needle. The voltage used was 35 kVwhich was high enough to initiate the spinning process with arcing.

The chitosan solution containing the drug was difficult, if notimpossible to spin. Adding the PVP hydrogel to the solution opened theway to spin the solution with some modifications to the percentage ofeach component of the solution. The drug concentration had no noticeableeffect on the spinability of the solution. Two trial solutions were spunsuccessfully with the following parameters: feeding solution at a rateof 2.5 mL/hour, traverse needle speed 20 cm/min, rotational of thecollector speed is 100 rpm, cabinet temperature 35° C., evacuation airflow rate 0.25 m³/min, positive voltage of 23 kV of DC, and negativevoltage −4 kV of DC. The composition of the first solution A comprised30 mL solution of 1.5% (wt/v) chitosan, 30 mL solution of 8% (wt/v) PVA,and 0.855 g ciprofloxacin; and the second solution B contained 25 mLsolution of 1.5% chitosan, 8 mL solution of 10% (wt/v) PVA, and 0.304 gciprofloxacin. Both solution A and B were prepared by dissolvingchitosan in 2% acetic acid and PVA and ciprofloxacin were dissolved indeionizer water.

The morphology of the deposited nanofibers on the medical mesh werecoated with 3 nanometer of gold layer then examined by Scan ElectronicMicroscope. Sample A (FIG. 2A) shows average fiber diameter of about 110nm and average bore size, i.e., the diameter of the opening betweeninterlacing fibers, of about 700 nm. Sample B (FIG. 2B) shows averagefiber diameter of about 220 nm and average bore size of about 500 nm.

Example 3 Drug Relase Determination of Ciprofloxacine-HCl CalibrationCurve

Column, Agilent Zorbax extend-C18 column (150 mm length×4.6 mm, i.d., 5μm), mobile phase, 0.025 M σ-phosphoric acid adjusted to pH 3 withtriethylamine:acetonitrile (75:25), UV detector set at λ=278 nm, flowrate of 1 mL/min, and injection volume of 20 μL. Approximately 20 mg ofciprofloxacin-HCl standard was weighed, transferred into a 100 mLvolumetric flask, dissolved in methanol and volume was completed withphosphate buffer pH 7.4. The stock standard solution (0.2 mg/mL) werediluted to give a concentration range of 1.6 to 40 mg/mL using phosphatebuffer pH 7.4 as diluting solvent.

In Vitro Drug Release Study

In vitro drug release studies from 2×2 cm² pieces of nanofiber mesh werecarried out over 12 days at specified times. The mesh of each formula ofTable 1 was placed in bottles with 20 mL PBS, pH 7.4. The bottles wereplaced in oscillating water bath at 37° C.±2. Aliquots of 1 mL werewithdrawn at different time intervals and replaced each time with freshPBS. The solution was filtered and 20 μL was injected to HPLC (Agilenttechnology, Germany) and area under peak was detected for each timeintervals to determine the concentrations of ciprofloxacin in each time.

Kinetic Analysis of In Vitro Drug Release Data

The in vitro release data were analyzed by zero-order, first-order, anddiffusion controlled mechanism according to the simplified Higuchi modeland the results are summarized in Table 2. The selection of a mechanismwas based on the determined regression coefficient (R²) for theparameters studied, wherein the order of release is determined bylargest regression coefficient [Dangprasirt et al. “Development ofdiclofinac sodium controlled release solid dispersion powders andcapsules by freeze drying technique using ethylcellulose and chitosan ascarriers” Drug Development and Industrial Pharmacy (1998) 24: 947-53].

TABLE 2 Regression coefficients of different mathematical fitted releaseof ciprofloxacin-HCl from different nanofibers formulations. Nanofibersformulations Zero order First order Higuchi A 0.832 0.774 0.928 B 0.7410.674 0.931 C 0.708 0.712 0.971

The Higuchi model displays the best fit to the data obtained for thedrug release from all formulas of nanofiber mats, because drug diffusionout of the nanofiber mats is the main factor in drug release. Theobserved result is in agreement with the results of Jannesari et al.[“Composite poly (vinyl alcocol)/Poly vinyl acetate elecrospunnanofiberus mats as a novel wound dressing matrix for controlled releaseof drugs. International journal of nanomedicine” (2011) 6: 993-1003,incorporated herein by reference in its entirety]. When a drug embeddedin a polymers come into contact with a liquid hydrate, a gel layer isformed at the surface of the polymer that is essential for sustainingand controlling drug release from polymer. FIG. 3 shows the convexrelease curves for formulation A, B, and C. Each curve has an initialburst release phase followed sustained release phase. There was nosignificant difference between the three formulations. The release ofthe drug from biodegradable polymer is governed by the combination ofboth mechanisms which depends on the relative rates of erosion anddiffusion. Most biodegradable polymers used for drug delivery aredegraded by hydrolysis. As water molecules break the chemical bondsalong the polymer chain, the physical integrity of the polymer degradesand allows drug to be released.

Example 4 Differential Scanning Colorimetry (DSC)

DSC (Netzch, Japan) thermograms for samples of ciprofloxacin-HCl, PVP,PVA and mixtures thereof were recorded and analyzed. Approximately 2 mgof samples were weighed into DSC aluminum pans and were crimped followedby heating under nitrogen flow (20 mL/min) at a heating rate 5° C./minfrom 25-300° C. Aluminum pan containing same quantity of indium was usedas reference. FIG. 4 shows the DSC thermograms spectra ofciprofloxacin-HCl, PVP, PVA, and their physical mixture.Ciprofloxacin-HCl shows a sharp endothermic melting peak at 145° C.,which was shifted 135° C. in the thermogram of the physical mixture withno appearance of new peak. Therefore, no incompatibility was observed.

Example 5 Evaluation of Antibacterial Activity

Antibacterial activity of ciprofloxacin alone against selected strainsof Staphylococcus aureus (ATCC 12228), Escherichia coli (ATCC 25922),and Pseudomonas aeruginosa (ATCC 27853) was carried out by thestandardized cup diffusion described by the Clinical and LaboratoryStandards Institute [CLSI “Performance standards for antimicrobialsusceptibility testing” 19 Ed. Informational Supplement, (2009) Vol. 29,Document M100-S19, Clinical Laboratory Standards Institute. Wayne, Pa.].

Preparation of Bacterial Inoculum:

Colonies from overnight cultures of the strains were suspended in salinesolution and appropriately diluted to match the turbidity standard of0.5 on McFarland Scale, diluted in saline (1:100) and used for seedinoculation of 25 mL agar plates or liquid medium to give the finalcount of 2-5×10⁵ CFU/mL as determined by viable count technique.

In Vitro Antimicrobial Test

The antimicrobial activity of the casting or electrospun nanofiberspolymer containing antibacterial agent was tested against the selectedmicroorganisms. The polymer was cut into 1 cm² discs. Antibiotic freepolymer was used as control. The tests of antibacterial activities werecarried out by one of two method as described by Xu et al. [“UltrafinePEG-PLA fibers loaded with both paclitaxel and doxorubicin HCl and theirin vitro cytotoxicity” European Journal of Pharmaceutics andBiopharmaceutics (2009) 72:18-25].

Disc Diffusion Method

The experiment was performed in Mueller Hinton agar (Oxoid, USA) plateusing a modified Kirby Bauer technique [Bauer et al. “Antibioticsusceptibility testing by standardized single disk method” Am J ClinPathol (1966) 45:493-496]. Inoculated agar plates were prepared and 1cm² discs of the polymer were placed on the surface of the seedinoculated agar plates, then the agar plates were incubated at 37° C.for 24 h. After incubation, the inhibition zones surrounding the samplewere measured.

Evaluation of the Activity of the Released Antibiotic from the Polymer(Cast or Nonofiber Mat)

The experiment was carried out by cup diffusion assay in which a pieceof 1 cm² disc of the polymer was immersed in 0.1 M 1 mL phosphatebuffer, pH 7.4. The phosphate buffer was replaced by another aliquotevery 24 hr for 20 days incubation period at 37° C. Each phosphatebuffer sample was used to evaluate antimicrobial activity of thereleased drug by cup diffusion assay. The formed zones of inhibitionwere measured. Discs of non-medicated polymer were used as blankcontrol.

Determination of Minimum Inhibitory Concentration (MIC) of CiprofloxacinAgainst Staphylococcus aureus and E. coli

Evaluation of the activity of CP against S. aureus and E. coli referencestrains was carried out by cup diffusion assay. According to theobtained results, the MICs of CP against S. aureus and E. coli standardstrains were 0.32 and 0.04 μg/ml, respectively.

The Antibacterial Activity of Ciprofloxacin in the Chitosan Polymer

Mueller Hinton agar plates inoculated with one of the testedmicroorganisms was prepared as described. A 1 cm² piece of medicated CHpolymer as casting film was placed on the agar surface and incubated at37° C. for 18-24 h. After incubation a clear zone of inhibition wasobserved in all the plates with range of size between 42 and 51 mm.Also, a piece of non-medicated CH polymer without CP was used asnegative control. A slight activity was observed with an inhibitionzones in the range of 17-21 mm. The results indicated that the formulaused in preparing the medicated casting film is suitable for releasingthe drug with concentration sufficient to inhibit the microorganisms inthe surrounding area.

The antimicrobial activity of chitosan comprising CP against strains ofE. coli and S. aureus was monitored over 14 days. The cultures wereincubated over 14 days from inoculation at 37° C. The E. coli culturesshowed inhibition zones formed and decreased 60 to 55 mm during thefirst 9 days from the rest of the incubation period and remainedconstant throughout the last five days. Similarly, the S. aureus culturedisplayed inhibition zone that decreased from 51 to 45 mm in the first 6days and remained constant thereafter until the end of 14 days (see FIG.5). In control experiment, the non-medicated chitosan casting filmwithout CP showed inhibition zones in the range 17-23 mm in the initial2-4 days and no antibacterial activity is observed thereafter for theremaining duration of the 14 days.

Antibacterial Activity of Ciprofloxacin in the Medicated ElectrospunNanofiber

A 1 cm² piece of the medicated nanofiber was placed on the surface ofMueller Hinton agar plates inoculated with one of the testedmicroorganisms and incubated at 37° C. for 18-24 h. After incubation, aclear zone of inhibition was observed in all the plates ranging in sizebetween 46 and 54 mm. The observed result indicated that the formulaused in preparing the medicated nanofiber is suitable for releasing ofthe CP with concentration sufficient to inhibit the microorganisms inthe surrounding area.

The antimicrobial activity of electrospun nanofiber mats of theinvention against strains of E. coli and S. aureus was monitored over 14days. For the E. coli cultures, a small reduction in the initiallyformed inhibition zone from 60 to 55 mm in the initial 8 days frominoculation followed by stable inhibition zone in the remaining periodof 14 days. Similar observation is observed in S. aureus straincultures. Gradual decrease in the inhibition zone from 51 to 45 mm isobserved in the first 9 days followed by stable inhibition zone in theremaining period of the 14 days (FIG. 6).

Example 6 Surgical Animal Model

Rabbits were used as animal model. They were anesthetized using theappropriate dose of Ketamine and Xylazine. Following hair clipping,sterile prep using povidone iodine solution, the scalpel was used tomake a 4 cm skin incision on the abdomen of the animal. The subcutaneoustissue was dissected and a pocket was created under the abdominalmuscles using a combination of sharp and blunt dissection. According tothe animal specific group, a standardized piece of either regularpolypropylene mesh or the newly created mesh was placed in thesubmuscular pocket along with the appropriate inoculum. The wounds wereclosed with interrupted Nylon sutures. The animals were followed 2 weeksprior to analysis. The rabbits were divided into Groups A-E and treatedas described below.

Group A animals n=4 (control animals) received a standard piece ofnon-medicated and medicated polypropylene mesh and 100 μL of sterilenormal saline into the surgical wound.

Group B animals n=4 received a standard piece of polypropylene mesh andare inoculated with one mL solution containing 10⁸ bacterial cells,referred herein as a standard dose, of Staphylococcus aureus into thewound.

Group C animals n=4 received a standard piece of theantibiotic-containing nanofiber mesh of the invention and are inoculatedwith a standard dose of Staphylococcus aureus into the wound.

Group D animals n=4 received a standard piece of polypropylene mesh andare inoculated with a standard dose of Escherichia coli into the wound.

Group E animals n=4 received a standard piece of theantibiotic-containing nanofiber mesh of the invention and are inoculatedwith a standard dose of Escherichia coli into the wound.

Each group followed for 2 weeks prior to euthanasia. At necropsy, themesh and surrounding tissue are excised under sterile conditions. Eachspecimen is placed in 10 mL of sterile 0.9% normal saline, homogenizedand submitted for in vitro testing. Briefly, 50 μL of homogenate wasadded to Muller-Hinton agar plates, mixed well, allowed to solidify,incubated at 37° C. for 24 hrs, and the number of colonies were counted(CFU/mL). Each experiment was done in triplicates (n=3) and the resultsare summarized in Table 3.

TABLE 3 CFU/mL in medicated and non-medicated meshes Non-medicatedmeshes Medicated meshes ^(a)Sample Count (CFU/mL) Sample Count (CFU/mL)1 (control) — 11 (control) — 2 (control) — 12 (control) — B3 (S. aureus)9480 C13 (S. aureus) — B4 (S. aureus) 9333 C14 (S. aureus) 210 B5 (S.aureus) 26866 C15 (S. aureus) 1990  B6 (S. aureus) 23866 C16 (S. aureus)230 D7 (E. coli) 5840 E17 (E. coli) 530 D8 (E. coli) 24800 E18 (E. coli)— D9 (E. coli) 2326 E19 (E. coli) — D10 (E. coli) 25566 E20 (E. coli)320 ^(a)The first letter of the designation refers to the animals groupwhich are treaded as described above followed by the animal number andthe bacteria used.

The antibacterial activity of ciprofloxacin released from polymer coatedelectrospun nanofiber mats for samples (11, 13 and 14) from 14 daysincubation period at 37° C. against E. coli showed a gradual decrease inthe antibiotic activity of CP along the third, tenth, and fifth dayrespectively, while against S. aureus showed a gradual decrease in theactivity of CP along the fifth, tenth, and tenth day respectively, as aresult, sample number 13 was chosen for in vivo study.

The microbial count after 14 days of animal inoculation with S. aureusand E. coli revealed a drastic decrease in case of medicated meshescompared to the non-medicated ones in a percentage of (100%, 97.75%,92.6% and 99%) in group C S. aureus inoculated meshes and (91%, 100%,100% and 98.75%) in group E E. coli inoculated meshes (Table 3, FIG. 8).

1-14. (canceled)
 15. A method of making an implantable medicalprosthetic mesh comprising nanofibers and a polymer mesh, said methodcomprising: mixing a first solution containing chitosan in an amount inthe range of 0.5 wt. % to 5 wt. % with a second solution comprising anantibiotic in an amount in the range of 1.0 wt. % to 5.0 wt. % and apolymer blend in an amount in the range of 6 wt. % to 15 wt. %, andelectrospinning the solution mixture to form the nanofibers, wherein thenanofibers comprise: the chitosan in an amount in the range of 8 wt. %to 35 wt. %, the chitosan having a degree of deacetylation in a range of70% to 95%; the antibiotic in an amount in the range of 15 wt. % to 35wt. %; and the polymer blend in an amount in the range of 45 wt. % to 70wt. %, each weight percent relative to a total weight of the nanofibers;wherein the polymer blend comprises polyvinyl alcohol andpolyvinylpyrrolidone; wherein an average diameter of the nanofibers isin a range of 50 nm to 300 nm; wherein the nanofibers have a bore sizein a range of 300 nm to 900 nm, then depositing the nanofibers onto amesh substrate to form the implantable medical prosthetic mesh havingthe nanofibers present on the surface of the mesh substrate; and whereinthe antibiotic is released from the nanofibers steadily for at least 14days, and wherein the implantable medical prosthetic mesh has anantibiotic release rate in a range of 0.01 μg/(cm²·min) to 10μg/(cm²·min).
 16. The method of claim 15, wherein the ratio of the firstsolution/the second solution is in the range of 1:5 to 1:1. 17.(canceled)
 18. The method of claim 15, wherein the antibiotic isciprofloxacin.
 19. (canceled)
 20. The method of claim 15, whereinelectrospinning is carried out by an upward electrospinning method. 21.(canceled)
 22. (canceled)
 23. The method of claim 15, wherein theantibiotic is a penicillin, tetracycline, cephalosporin, quinolone,lincomycin, macrolide, sulfonamide, glycopeptide, aminoglycoside,carbapenem, or combination thereof.
 24. The method of claim 15, whereinthe antibiotic is a quinolone.
 25. The method of claim 15, wherein theantibiotic is selected from the group consisting of lomefloxacine,ofloxacin, gatifloxacin, norfloxacin, ciprofloxacin, moxifloxacin,levofloxacin, gemifloxacin, delafloxacin, cinoxacin, nalidixic acid,trovafloxacin, sparfloxacin, and combination thereof.
 26. The method ofclaim 15, wherein the polymer blend consists of polyvinyl alcohol andpolyvinylpyrrolidone.
 27. The method of claim 15, wherein the meshsubstrate comprises at least one of polypropylene,polytetrafluroethylene, polyethylene terephthalate, and polyvinylidenefluoride.
 28. The method of claim 15, wherein the mesh is at least oneof polyester, polysaccharide, and polyurethane.
 29. The method of claim15, wherein the nanofibers consist of chitosan, the polymer blend, andthe antibiotic.